A. Magnetic Resonance Imaging
Magnetic resonance imaging (MRI) refers generally to a form of clinical imaging based upon the principles of nuclear magnetic resonance (NMR). Any nucleus which possesses a magnetic moment will attempt to align itself with the direction of a magnetic field, the quantum alignment being dependent, among other things, upon the strength of the magnetic field and the magnetic moment. In MRI, a uniform magnetic field B.sub.0 is applied to an object to be imaged; hence creating a net alignment of the object's nuclei possessing magnetic moments. If the static field B.sub.0 is designated as aligned with the z axis of a Cartesian coordinate system, the origin of which is approximately centered within the imaged object, the nuclei which posses magnetic moments precess about the z-axis at their Larmor frequencies according to their gyromagnetic ratio and the strength of the magnetic field.
Water, because of its relative abundance in biological tissues and its relatively strong net magnetic moment M.sub.z created when placed within a strong magnetic field, is of principle concern in MR imaging. Subjecting human tissues to a uniform magnetic field will create such a net magnetic moment from the typically random order of nuclear precession about the z-axis. In a MR imaging sequence, a radio frequency (RF) excitation signal, centered at the Larmor frequency, irradiates the tissue with a vector polarization which is orthogonal to the polarization of B.sub.0. Continuing our Cartesian coordinate example, the static field is labeled B.sub.a while the perpendicular excitation field B.sub.1 is labeled B.sub.xy. B.sub.xy is of sufficient amplitude and duration in time, or of sufficient power to nutate (or tip) the net magnetic moment into the transverse (x-y) plane giving rise to M.sub.xy. This transverse magnetic moment begins to collapse and re-align with the static magnetic field immediately after termination of the excitation field B.sub.1. Energy gained during the excitation cycle is lost by the nuclei as they re-align themselves with B.sub.0 during the collapse of the rotating transverse magnetic moment M.sub.xy.
The energy is propagated as an electromagnetic wave which induces a sinusoidal signal voltage across discontinuities in closed-loop receiving coils. This represents the NMR signal which is sensed by the RF coil and recorded by the MRI system. A slice image is derived from the reconstruction of these spatially-encoded signals using well known digital image processing techniques.
B. Local Coils and Arrays
The diagnostic quality or resolution of the image is dependent, in part, upon the sensitivity and homogeneity of the receiving coil to the weak NMR signal. RF coils, described as "local coils" may be described as resonant antennas, in part, because of their property of signal sensitivity being inversely related to the distance from the source. For this reason, it is important to place the coils as close to the anatomical region-of-interest (ROI) as possible.
Whereas "whole body" MRI scanners are sufficiently large to receive and image any portion of the entire human body, local coils are smaller and therefore electromagnetically couple to less tissue. Coupling to less tissue gives rise to coupling to less "noise" or unwanted biologically or thermally generated random signals which superimpose upon the desired MR signal. The local coils may be of higher quality factor (Q) than the body coils due to their smaller size. For all of these reasons, local coils typically yield a higher signal-to-noise S/N ratio ratio than that obtainable using the larger whole body antenna. The larger antenna is commonly used to produce the highly homogenous or uniform excitation field throughout the ROI, whereas the local coil is placed near the immediate area of interest to receive the NMR signal. The importance of accurate positioning leads to the development of local coils which conform to the anatomy of interest, yet function to permit ease of use.
While the smaller local coil's size works to an advantage in obtaining a higher S/N ratio, this reduced size also presents a disadvantage for imaging deep-seated tissues. Typically, the single- conductor coil diameter which yields the optimal S/N ratio at a depth `d` is a coil of diameter `d` ; hence, larger diameter single-conductor coils are required to image regions in the abdomen and chest of human patients. This increased coil size results in less than desirable performance, both in terms of S/N ratio and homogeneity of the sensitivity profile (which effects the uniform brightness of the image), and offers little advantage over the body coil of the system.
The S/N ratio of the NMR signal may be further increased by orienting two coils, or coil pairs about the imaged object so that each detects RF energy along one of a pair of mutually perpendicular axes. This technique is generally known as quadrature detection and the signals collected are termed quadrature signals.
The outputs of the quadrature coils are combined so as to increase the strength of the received signal according to the simple sum of the output signals from the coils. The strength of the noise component of these signals, however, will increase only according to the square root of the sum of the squares of the uncorrelated noise components. As a result, the net S/N ratio of the combined quadrature signals increases by approximately .sqroot.2 over the S/N ratio of the individual coils.
The quadrature orientation of the two coils introduces a 90.degree. phase difference between the NMR signals detected by these coils. Therefore, combining the outputs from the two quadrature coils to achieve the above described signal-m-noise ratio improvements requires that one signal be shifted to have the same phase as the other signal so that the amplitudes of the signals simply add in phase.
The approximate net gain of .sqroot.2 in S/N ratio is achievable primarily due to the lack of inductive coupling between the coil pairs. This ensures that only the uncorrelated noise components add, in lieu of both the uncorrelated and correlated noise components, to reduce the effective S/N ratio. Inductive isolation is achieved by geometrically orienting the coil conductors such that the mutual inductance is minimized between the coil pairs according to the following: ##EQU1## where M represents the mutual inductance between coils 1 and 2 and the vector components dl.sub.1, and dl.sub.2 represent segments of coils 1 and 2 with current amplitudes I.sub.1 and I.sub.2. The denominator represents the magnitude difference of the position vectors of each dl segment. The condition wherein M is approximately zero with respect to the individual self inductances of coils 1 and 2, is known as inductive isolation between the coils.
A method of increasing the S/N ratio of the NMR signal over a larger region is to digitally add the post processed signals derived from more than one coil; each sensitive to the precessing nuclei within overlapping volumes. If two coils' signals are processed and converted into image data separately and then added digitally, one can obtain an increase in S/N ratio within the larger volume. Separate amplifiers, analog-to-digital converters, and image processor channels represent an alternative configuration for processing the two signals in lieu of a single quadrature combiner. A system of four channels whose signals are derived from an array of four coils is described in U.S. Pat. No. 4,825,162. In the '162 patent, an array of coils is described wherein the adjacent coils overlap to prevent nearest-neighbor interaction (inductive coupling). The interaction between the next-nearest-neighbor is supposedly reduced by connection of each coil of the array to low input impedance preamplifiers.
The problem with this solution is, among other things, the use of preamplifiers with low input impedance. This additional circuitry is costly and adds another set of possible failure modes into the system. This preamplifier circuitry is sensitive to coil impedance changes resulting from patient loading variations as well as to noise spikes or power surges within the receiver chain.
One can minimize the effects of next-nearest-neighbor interaction if one properly utilizes the formulation above and in the following arguments to minimize inductive coupling between all resonant structures. In this case, the additional preamplifier circuitry is no longer required.
First, nearest-neighbor or adjacent coil interaction is a much more dominant coupling than the next-nearest-neighbor coupling--usually one or two orders of magnitude larger depending upon coil size and spacing. Second, if near-neighbor coupling has not been sufficiently minimized, then next-nearest neighbor coupling will occur via neighbor-to-neighbor interaction as strongly as, or stronger than inductive coupling between next-nearest-neighbors. Third, next-nearest neighbor interaction (inductive coupling) is further reduced towards zero when the next-nearest-neighbor coils are dominantly loaded by coupling to patient tissues. Such is the case in mid to high field scanners operating above 20 MHz. The coil's impedance is also dominated resulting from coupling to eddy current loops generated within the patient tissues. This is predicted from the mutual impedance formulation ##EQU2## where Z.sub.1d is the driving or output impedance of coil 1, Z.sub.11 is the self-impedance of coil 1, (I.sub.2 /I.sub.1) is the ratio of induced eddy currents (loop 2) to the current in coil 1, and Z.sub.12 is the mutual impedance between the loops which is equal to the radian frequency times the mutual inductance between said loops.
The implication from the above three facts is as follows. If one ensures consistent and dominant loading of the coil elements and if one ensures that near-neighbor coupling has been minimized (that is, inductive isolation has been achieved), and if the antenna element size, geometrical orientation, and spacing is designed so as to minimize next-nearest neighbor coupling, then the array will work properly with little degrading interaction amongst the elements.
Inductive isolation is achieved by geometrically orienting two coil conductors such that their mutual inductance is minimized according to the above formulation. The condition wherein M is approximately zero with respect to the individual self inductances of coils 1 and 2, is known as geometric isolation between the coils. This is a special case of inductive isolation but is restrictive in application, as discussed below.
As the coil geometries are sufficiently large or close to the surrounding system conductors (antenna, faraday screen, cryostat tubing, etc.) in addition to the biological conducting medium, this coupling formula must be extended to include M=M.sub.12 +M.sub.13 +M.sub.23 ; where the first term on the right side of the equation is as described above, and the latter two terms define the coupling resulting from each coil's coupling to eddy current loops (loop 3) generated on or within the surrounding conductors (system or biological). These additional coupling terms must be accounted for in the adjustment of conductor geometries with respect to each other spatially. With these terms taken into account, the proper critical spacing may be found between coil loops 1 and 2.
Phased array coils such as described in U.S. Pat. Nos. 4,825,162 and 5,198,768 only utilize linear coil technology to create an array of coils; each sensitive to one vector component of the NMR signal in a unique imaging volume to create a coil set with a large ROS. Both of those patents focus on obtaining a larger region of sensitivity using a bank of coils whose signals input to separate preamplifiers and digital reconstruction ports on a computer. This does not produce the optimal S/N ratio from within the imaging volumes such as would be the case where two coil elements are sensitive to orthogonal vector components of magnetic flux within each volume (i.e. an array of quadrature elements). The prior art is reliant upon geometric isolation and/or low impedance preamps with no compensation for eddy current-induced coupling. The prior art also restricts the geometry to a planar surface only. This restriction is due, in part, to the fact that the prior art is dependent upon geometric isolation only, and this alone is inadequate to ensure sufficient isolation between non-adjacent pairs of non-planar conductors. This type of coil array presents engineering challenges in maintaining isolation between elements as additional elements are co-located in space, thereby increasing the potential for inductive coupling.